Magnetic Resonance Imaging (MRI) utilizes hydrogen nuclear spins of the water molecules in the human body, which are polarized by a strong uniform static magnetic field, the B0 field. The magnetically polarized hydrogen nuclear spins generate magnetic moments and precess in the direction of the B0 field and produce no useful information unless disturbed from the equilibrium state by an excitation.
The generation of a nuclear magnetic resonance (NMR) signal for MRI data acquisition is accomplished by exciting the magnetic moments with a uniform radio-frequency (RF) magnetic field, the B1 field, applied transverse to the B0 field. This B1 field is centred at the precessional frequency of the protons (Larmor frequency) and causes some of the protons to change their spin direction by some predetermined angle. The B1 field is produced by an RF transmit coil that is driven by a computer-controlled RF transmitter with a RF power amplifier. The application of the B1 field has the effect of nutating the net magnetization and at the same time causes the magnetic moments to gain magnetic energy from the applied B1 field. After the application of the B1 field ceases, the magnetic moments revert to their ground state (through a process of free induction decay) and in doing so induce a measurable MR signal in a receiver RF coil that is tuned to the Larmor frequency. The receive RF coil can either be the transmit coil itself or an independent receive-only RF coil. The detected MR signal is processed to produce MR images by using additional pulsed magnetic gradient fields that are generated by gradient coils integrated inside the main magnet system. The gradient fields are used to spatially encode the signals and selectively excite a specific volume of the human body. There are usually three sets of gradient coils in a standard MRI system that generate magnetic fields in the same direction as the main magnetic field and vary linearly in the imaging volume.
In MRI, it is desirable for the excitation of the B1 field and reception of the MR signal to be spatially uniform in the imaging volume for high quality MR images. In a standard MRI system, the transmission of the B1 field is generally through the MRI system whole-body volume RF coil. This whole body RF coil, however, produces lower signal-to-noise ratio (SNR) if it is also used for the reception of the MR signal, mainly because of the large distance from the volume under imaging to the coil itself. Therefore, in order to achieve a high SNR, special-purpose RF coils are used for receiving the MR signal. In practice, a well-designed specialty RF coil has the following functional properties: high SNR, highly uniform sensitivity, high unloaded quality factor (Q) of the resonance circuit, and high ratio of the unloaded to loaded Q factors. In addition, the RF coil device must be mechanically designed to facilitate patient handling, comfort and safety. Improvement in the SNR of the detected MR signal can be achieved by using a small local coil placed close to the human body. As this local coil is placed close to the proximity of the region of interest, the small reception pattern using this local coil can focus in the region of interest thus improving the SNR. An array of these small local coils can be used to increase the coverage of the region of interest and this array system is generally referred to as phased array RF coils (see for example U.S. Pat. No. 4,825,162 assigned to General Electric Company). The outputs from the phased array system are simultaneously processed and the MR images are combined using a sum-of-square method. The phased array system obtains the high SNR and resolution of a small local coil over a large field-of-view (FOV) normally associated with body imaging but with no increase in imaging time.
In our co-pending international patent application number PCT/AU2006/000311, a focusing scheme is described for a phased array coil system that further increases the quality of image obtained. The invention is described with reference to a number of small local coils with particular application to the head and chest. The content of the co-pending application is incorporated herein by reference.
Phased array coil structures usually display strong mutual coupling between individual coil elements and some of the undesirable effects include difficulty in tuning, reduced SNR and RF field distortion causing image artefacts. Hence, minimizing the mutual coupling is known to be important to the quality of the images produced.
A number of methods have been suggested to minimize mutual coupling. Some of the known methods include the overlapping of adjacent coils (U.S. Pat. No. 4,825,162), the use of a magnetic decoupling circuit (U. S. patent application Ser. No. 2005/0275403), a degenerate birdcage coil design (U.S. Pat. No. 7,180,291), employing capacitive decoupling networks (see for example U.S. Pat. No. 7,091,721 assigned to IGC-Medical Advances Inc) and the use of low input impedance pre-amplifiers.
Another document which generally discloses the field of the invention is Japanese patent number 08-187235, assigned to GE Yokogawa Medical Syst Ltd. This patent discloses a birdcage coil for MRI having a number of diode-and-inductor in series circuits connected in parallel to capacitors in one ring of the birdcage coil so as to decouple the birdcage coil from another coil. This patent does not relate to coupled counter-wound inductors for decoupling coil elements.
There are some constraints, however, in using these decoupling methods. The overlapping of adjacent coils sacrifices the area of coverage, lumped-element decoupling networks have limitations on their decoupling power and the use of low input impedance preamplifiers can limit power transfer and limit the use of the phased array coils to receive only (i.e not suitable for transceive operation).